The field of the invention is coherent imaging using vibratory energy, such as ultrasound and, in particular, ultrasound imaging of the heart.
There are a number of modes in which ultrasound can be used to produce images of objects. The ultrasound transmitter may be placed on one side of the object and the sound transmitted through the object to the ultrasound receiver placed on the other side (“transmission mode”). With transmission mode methods, an image may be produced in which the brightness of each pixel is a function of the amplitude of the ultrasound that reaches the receiver (“attenuation” mode), or the brightness of each pixel is a function of the time required for the sound to reach the receiver (“time-of-flight” or “speed of sound” mode). In the alternative, the receiver may be positioned on the same side of the object as the transmitter and an image may be produced in which the brightness of each pixel is a function of the amplitude or time-of-flight of the ultrasound reflected from the object back to the receiver (“refraction”, “backscatter” or “echo” mode). The present invention relates to a backscatter method for producing ultrasound images.
There are a number of well known backscatter methods for acquiring ultrasound data. In the so-called “A-scan” method, an ultrasound pulse is directed into the object by the transducer and the amplitude of the reflected sound is recorded over a period of time. The amplitude of the echo signal is proportional to the scattering strength of the refractors in the object, and the time delay is proportional to the range of the refractors from the transducer. In the so-called “B-scan” method, the transducer transmits a series of ultrasonic pulses as it is scanned across the object along a single axis of motion. The resulting echo signals are recorded as with the A-scan method and either their amplitude or time delay is used to modulate the brightness of pixels on a display. With the B-scan method, enough data are acquired from which an image of the refractors can be reconstructed. The so-called “M-scan” is very similar to the B-scan in that it is a continuous series of B-scans. This mode is commonly used to show motion of the heart so that heart structures can be observed during all phases of the cardiac cycle.
Ultrasonic transducers for medical applications are constructed from one or more piezoelectric elements sandwiched between a pair of electrodes. When used for ultrasound imaging, the transducer typically has a number of piezoelectric elements arranged in an array and driven with separate voltages (apodizing). By controlling the time delay (or phase) and amplitude of the applied voltages, the ultrasonic waves produced by the piezoelectric elements (transmission mode) combine to produce a net ultrasonic wave focused at a selected point. By controlling the time delay and amplitude of the applied voltages, this focal point can be moved in a plane to scan the subject.
The same principles apply when the transducer is employed to receive the reflected sound (receiver mode). That is, the voltages produced at the transducer elements in the array are summed together such that the net signal is indicative of the sound reflected from a single focal point in the subject. As with the transmission mode, this focused reception of the ultrasonic energy is achieved by imparting separate time delay (and/or phase shifts) and gains to the signal from each transducer array element.
This form of ultrasonic imaging is referred to as “phased array sector scanning”, or “PASS”. Such a scan is comprised of a series of measurements in which the steered ultrasonic wave is transmitted, the system switches to receive mode after a short time interval, and the reflected ultrasonic wave is received and stored. Typically, the transmission and reception are steered in the same direction (θ) during each measurement to acquire data from a series of points along a scan line. The receiver is dynamically focused at a succession of ranges (R) along the scan line as the reflected ultrasonic waves are received. The time required to conduct the entire scan is a function of the time required to make each measurement and the number of measurements required to cover the entire region of interest at the desired resolution and signal-to-noise ratio.
A primary problem in ultrasonic imaging has been that many of the body's internal structures have similar characteristics as regards the reflection of ultrasonic energy, so that it is difficult to obtain as clear and detailed images of many of the structures as is desired. In particular, many of the structures of interest, such as the muscles of the heart, are perfused with blood, so that it is difficult to distinguish between blood vessels and the chambers of the heart and the heart muscles.
One solution to this problem has been ultrasonic imaging using contrast agents injected into the blood stream. Ultrasonic contrast agents are now commercially available and are essentially small bubbles of gas, such as air, formed by agitating a liquid or bubbling gas through a liquid, such as a saline solution or a solution containing a bubble forming compound, such as albumin. When insonicated, the bubbles resonate at their resonant frequency and emit energy at both the fundamental and second harmonic of their resonant frequency, thereby returning an enhanced signal at or around these frequencies and thereby providing an enhanced image of the liquid or tissue containing the contrast agent. It is also well known that the bubbles “disappear” when insonicated and the current theory is that the insonication ruptures the bubble's shell, thereby allowing the gas to dissipate into the surrounding liquid or tissue.
The use of ultrasonic contrast agents is thereby advantageous in allowing enhanced imaging using ultrasonics rather than x-rays, thereby eliminating the radiation hazard and allowing the use of equipment that is significantly less expensive and hazardous to use. Also, the agents are non-toxic and dissolve relatively quickly into waste products, such as air and albumin, that are normally found in the body and that are themselves non-toxic. Further, the insonication of the agent in itself destroys the agent, so that the agent can effectively be “erased” during the imaging process to a degree.
The left ventricle (LV) in a mammalian heart carries out the functions of suction and ejection, transiting functionally through short-lived phases known as isovolumic intervals. Ventricular disease or disturbed myocardial electrical activation primarily prolongs the isovolumic intervals, with either no significant change or a shortening of ejection and filling times.
At the cellular level, the isovolumic intervals are associated with active fluxes in myoplasmic and sarcolemmal calcium that either initiate or reverse interactions between cardiac myofilaments. At the tissue level, isovolumic intervals are associated with asynchronous but synergistic movements of the subendocardial and subepicardial regions. During isovolumic contraction (IVC), the subendocardial fibers that form a right-handed helix shorten, while the left-handed helically oriented subepicardial fibers lengthen simultaneously. Conversely, during isovolumic relaxation (IVR), the subepicardial fibers that form the left-handed helix lengthen, while the right-handed helically directed subendocardial fibers shorten briefly. An initial asymmetric deformation of the LV may represent a “flow-directing feature” of the myocardial wall mechanics that reverses the direction of blood flow.
Conventionally, the timing of mitral valve closure has been used for dividing the preejection period into 2 component intervals. The first component, also referred to as “electromechanical delay,” is in continuity with end-diastole and refers to the interval from the onset of the Q wave on surface electrocardiography to mitral valve closure. Isovolumic contraction is the period that follows mitral valve closure and is characterized by a rapid rise in LV pressure before opening of the aortic valve. Recent observations indicate that cardiac muscle shortening is initiated significantly before closure of the mitral valve. Whether blood flow during this early stage accelerates into an axial momentum that ultimately causes mitral valve closure remains unclear. This necessitates further clarification on the rheologic features and nature of LV performance during the preejection period.
Previous investigations have used echocardiography and magnetic resonance velocity mapping of blood circulation for deciphering the features of LV intracavitary blood flow patterns. An intriguing aspect of this flow is the occurrence of intracavitary vortices. Findings from in vitro experiments suggested that strong vortices are required at the onset of ventricular contraction near the LV outflow, and that without these vortices, the mitral valve would remain open at the onset of ventricular contraction. However, features of intracavitary flow visualized directly in vivo during isovolumic periods have remained a mystery. Time-resolved flow vector quantifications could provide insights into why the initial LV deformations are asymmetric.